Superconductor magnetic resonance imaging system and method (super-MRI)

ABSTRACT

Methods and apparatuses for magnetic resonance imaging (MRI) and/or magnetic resonance spectroscopy comprising a superconducting main magnet operable to generate a uniform magnetic field in an examination region, at least one superconducting gradient field coil operable to apply a respective at least one magnetic field gradient within the examination region, and at least one RF coil that is operable to transmit and receive radio frequency signals to and from the examination region, and that is configured for cooling and comprises at least one of (i) a non-superconducting material that when cooled to a temperature below room temperature has a conductivity higher than that of copper at that temperature and (ii) a superconducting material. The main magnet, the gradient coils, and each of the at least one RF coil of a given system may each be implemented as high temperature superconductor materials.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.12/416,606, filed Apr. 1, 2009, which claims the benefit of U.S.Provisional Application No. 61/159,008, filed Mar. 10, 2009, each ofwhich is hereby incorporated by reference herein in its entirety.

TECHNICAL FIELD

The present invention relates generally to magnetic resonance imagingand spectroscopy, and, more particularly, to magnetic resonance imagingand spectroscopy apparatus employing superconductor components, and tomethods for manufacturing such apparatus.

BACKGROUND

Magnetic Resonance Imaging (MRI) technology is commonly used today inlarger medical institutions worldwide, and has led to significant andunique benefits in the practice of medicine. While MRI has beendeveloped as a well-established diagnostic tool for imaging structureand anatomy, it has also been developed for imaging functionalactivities and other biophysical and biochemical characteristics orprocesses (e.g., blood flow, metabolites/metabolism, diffusion), some ofthese magnetic resonance (MR) imaging techniques being known asfunctional MRI, spectroscopic MRI or Magnetic Resonance SpectroscopicImaging (MRSI), diffusion weighted imaging (DWI), and diffusion tensorimaging (DTI). These magnetic resonance imaging techniques have broadclinical and research applications in addition to their medicaldiagnostic value for identifying and assessing pathology and determiningthe state of health of the tissue examined.

During a typical MRI examination, a patient's body (or a sample object)is placed within the examination region and is supported by a patientsupport in an MRI scanner where a substantially constant and uniformprimary (main) magnetic field is provided by a primary (main) magnet.The magnetic field aligns the nuclear magnetization of precessing atomssuch as hydrogen (protons) in the body. A gradient coil assembly withinthe magnet creates a small variation of the magnetic field in a givenlocation, thus providing resonance frequency encoding in the imagingregion. A radio frequency (RF) coil is selectively driven under computercontrol according to a pulse sequence to generate in the patient atemporary oscillating transverse magnetization signal that is detectedby the RF coil and that, by computer processing, may be mapped tospatially localized regions of the patient, thus providing an image ofthe region-of-interest under examination.

In a common MRI configuration, the static main magnetic field istypically produced by a solenoid magnet apparatus, and a patientplatform is disposed in the cylindrical space bounded by the solenoidwindings (i.e. the main magnet bore). The windings of the main field aretypically implemented as a low temperature superconductor (LTS)material, and are super-cooled with liquid helium in order to reduceresistance, and, therefore, to minimize the amount of heat generated andthe amount of power necessary to create and maintain the main field. Themajority of existing LTS superconducting MRI magnets are made of aniobium-titanium (NbTi) and/or Nb₃Sn material which is cooled with acryostat to a temperature of 4.2 K.

As is known to those skilled in the art, the magnetic field gradientcoils generally are configured to selectively provide linear magneticfield gradients along each of three principal Cartesian axes in space(one of these axes being the direction of the main magnetic field), sothat the magnitude of the magnetic field varies with location inside theexamination region, and characteristics of the magnetic resonancesignals from different locations within the region of interest, such asthe frequency and phase of the signals, are encoded according toposition within the region (thus providing for spatial localization).Typically, the gradient fields are created by current passing throughcoiled saddle or solenoid windings, which are affixed to cylindersconcentric with and fitted within a larger cylinder containing thewindings of the main magnetic field. Unlike the main magnetic field, thecoils used to create the gradient fields typically are common roomtemperature copper windings. The gradient strength and field linearityare of fundamental importance both to the accuracy of the details of theimage produced and to the information on tissue chemistry (e.g., inMRSI).

Since MRI's inception, there has been a relentless pursuit for improvingMRI quality and capabilities, such as by providing higher spatialresolution, higher spectral resolution (e.g., for MRSI), highercontrast, and faster acquisition speed. For example, increased imaging(acquisition) speed is desired to minimize imaging blurring caused bytemporal variations in the imaged region during image acquisition, suchas variations due to patient movement, natural anatomical and/orfunctional movements (e.g., heart beat, respiration, blood flow), and/ornatural biochemical variations (e.g., caused by metabolism during MRSI).Similarly, for example, because in spectroscopic MRI the pulse sequencefor acquiring data encodes spectral information in addition to spatialinformation, minimizing the time required for acquiring sufficientspectral and spatial information to provide desired spectral resolutionand spatial localization is particularly important for improving theclinical practicality and utility of spectroscopic MRI.

Several factors contribute to better MRI image quality in terms of highcontrast, resolution, and acquisition speed. An important parameterimpacting image quality and acquisition speed is the signal-to-noiseratio (SNR). Increasing SNR by increasing the signal before thepreamplifier of the MRI system is important in terms of increasing thequality of the image. One way to improve SNR is to increase the magneticfield strength of the magnet as the SNR is proportional to the magnitudeof the magnetic field. In clinical applications, however, MRI has aceiling on the field strength of the magnet (the US FDA's currentceiling is 3 T (Tesla)). Other ways of improving the SNR involve, wherepossible, reducing sample noise by reducing the field-of-view (wherepossible), decreasing the distance between the sample and the RF coils,and/or reducing RF coil noise.

Despite the relentless efforts and many advancements for improving MRI,there is nevertheless a continuing need for yet further improvements inMRI, such as for providing greater contrast, improved SNR, higheracquisition speeds, higher spatial and temporal resolution, and/orhigher spectral resolution.

Additionally, a significant factor affecting further use of MRItechnology is the high cost associated with high magnetic field systems,both for purchase and maintenance. Thus, it would be advantageous toprovide a high quality MRI imaging system that is capable of beingmanufactured and/or maintained at reasonable cost, permitting MRItechnology to be more widely used.

SUMMARY OF INVENTION

Various embodiments of the present invention provide methods andapparatuses for magnetic resonance imaging (MRI) and/or magneticresonance spectroscopy comprising: a superconducting main magnetoperable to generate a uniform magnetic field in an examination region;at least one superconducting gradient field coil operable to apply arespective at least one magnetic field gradient within the examinationregion; and at least one RF coil that is operable to transmit andreceive radio frequency signals to and from the examination region, andthat is configured for cooling and comprises at least one of (i) anon-superconducting material that when cooled to a temperature belowroom temperature has a conductivity higher than that of copper at saidtemperature, and (ii) a superconducting material.

In accordance with some embodiments of the present invention, the mainmagnet, the gradient coils, and the RF coil are each implemented assuperconductors using high temperature superconductor materials. Inalternative embodiments, the superconducting main magnet, and/or one ormore of the at least one superconducting gradient field coil, and/or theRF coil are all formed from a low temperature superconducting material.

In accordance with some aspects of the present invention, the at leastone gradient coil and the at least one RF coil are disposed in at leastone vacuum chamber having at least one non-magnetic and non-metallicwall disposed between the examination region and the gradient coil andthe at least one RF coil. Additionally, the at least one gradient coiland the at least on RF coil may be disposed in a common vacuum chambercomprising said at least one non-magnetic and non-metallic wall. Afurther vacuum chamber may be disposed between the common vacuum chamberand the examination region, wherein the further vacuum chamber comprises(i) a first wall formed from the at least one non-magnetic andnon-metallic wall of the common vacuum chamber, and (ii) a secondnon-magnetic and non-metallic wall spaced away from said first wall.

In accordance with some aspects of the present invention, the mainmagnet may be disposed in a first vacuum chamber, and the at least oneRF coil and the at least one gradient coil may be disposed in a secondvacuum chamber. Alternatively, in some embodiments, the main magnet, theat least one RF coil, and the at least one gradient coil may be disposedin respective vacuum chambers.

In accordance with various aspects of the present invention, the atleast one RF coil may be implemented as a two-dimensional electron gasstructure and/or as a carbon nanotube structure. In some embodiments,the at least one RF coil may comprise a coil array.

In accordance with various embodiments, one or more cooling systems maybe used for cooling the main magnet, the at least one gradient coil, andthe at least one RF coil. In some embodiments, the superconducting mainmagnet is configured for cooling by a first cryogenic cooling system,the at least one RF coil is configured for cooling by a second cryogeniccooling system, and the at least one gradient coil is configured forcooling by a third cryogenic cooling system. In some embodiments, thesuperconducting main magnet is configured for cooling by a firstcryogenic cooling system, and the at least one RF coil and the at leastone gradient coil are configured for cooling by a second cryogeniccooling system. In some embodiments, the superconducting main magnet,the at least one RF coil, and the at least one gradient coil areconfigured for cooling by a common cryogenic cooling system.

In accordance with some aspects of the present invention, the at leastone superconducting gradient field coil comprises three superconductinggradient field coils that are configured to provide magnetic fieldgradients in three respective orthogonal directions, one of thedirections being along the direction of the uniform magnetic field inthe examination region.

In accordance with various aspects of the present invention, a methodfor magnetic resonance imaging comprises applying a uniform magneticfiled in an examination region using a superconducting main magnet,applying at least one magnetic field gradient within the examinationregion using at least one respective superconducting gradient fieldcoil, and transmitting and receiving radio frequency signals to and fromthe examination region using at least one RF coil that is configured forcooling and comprises at least one of (i) a non-superconducting materialthat when cooled to a temperature below room temperature has aconductivity higher than that of copper at that temperature and (ii) asuperconducting material. The superconducting main magnet, each of theat least one superconducting gradient field coil, and each of the atleast one superconducting RF coil may all be formed from an HTSmaterial. The at least one superconducting gradient field coil maycomprise three superconducting gradient field coils that are configuredto provide magnetic field gradient in three orthogonal directions, oneof the directions being along the direction of the uniform magneticfield in the examination region.

It will be appreciated by those skilled in the art that the foregoingbrief description and the following detailed description are exemplaryand explanatory of the present invention, but are not intended to berestrictive thereof or limiting of the advantages which can be achievedby this invention. Additionally, it is understood that the foregoingsummary of the invention is representative of some embodiments of theinvention, and is neither representative nor inclusive of all subjectmatter and embodiments within the scope of the present invention. Thus,the accompanying drawings, referred to herein and constituting a parthereof, illustrate embodiments of this invention, and, together with thedetailed description, serve to explain principles of this invention.

BRIEF DESCRIPTION OF THE DRAWINGS

Aspects, features, and advantages of embodiments of the invention, bothas to structure and operation, will be understood and will become morereadily apparent when the invention is considered in the light of thefollowing description made in conjunction with the accompanyingdrawings, in which like reference numerals designate the same or similarparts throughout the various figures, and wherein:

FIG. 1A schematically depicts a schematic cross-sectional view of anillustrative superconductor MRI system, in accordance with an embodimentof the present invention;

FIG. 1B schematically depicts in more detail the upper cross-sectionalportion of the main magnet system shown in FIG. 1A, in accordance withan embodiment of the present invention;

FIG. 2A schematically depicts in more detail an oblique view of thegradient coil configuration of the illustrative superconductor MRIsystem of FIG. 1A, in accordance with an embodiment of the presentinvention;

FIG. 2B schematically illustrates a cylindrical x-oriented gradient coilof FIG. 2A depicted in a plan view, in accordance with an embodiment ofthe present invention;

FIGS. 3A-3D schematically depict different examples of coolingconfigurations that may be used within an MRI system according tovarious embodiments of the present invention; and

FIGS. 4A and 4B illustrate cross sectional views of an illustrative coilconfiguration associated with a superconducting MRI system employing acylindrical, solenoid main magnet structure, in accordance with someembodiments of the present invention.

DESCRIPTION OF EMBODIMENTS OF THE INVENTION

As will be understood by those skilled in the art, while the ensuingdescription is set forth in the context of an MRI system that may beused for examining a patient, embodiments of the present inventioninclude systems and methods for magnetic resonance spectroscopy.Additionally, as used herein, MRI includes and embraces magneticresonance spectroscopic imaging.

FIG. 1A schematically depicts a schematic cross-sectional view of anillustrative superconductor MRI system 100 in accordance with anembodiment of the present invention. Superconductor MRI system 100includes an examination region 180; a movable patient bed 190; amagnet/coil housing 130; a main magnet system (shown in more detail inFIG. 1B) comprising (i) a main magnet that includes superconductingcoils 104, superconducting correction coils 106, and a coil frame 108,(ii) thermal sink 110, (iii) cryogen container 112, (iv) thermal shield114, (v) main magnet vacuum chamber housing 116, and (vi) cryogenicsystem 160. The movable patient bed 190 can be slid in and out of theexamination region. At least the portion of the patient bed 190 which islocated in the main magnetic field is made of non-metallic andnon-magnetic material such as plastic.

In the embodiment of FIGS. 1A and 1B, the superconductor main magnetsystem is implemented as a solenoid magnet that generates asubstantially uniform, horizontal magnetic field in the range of, forexample, about 0.5 T (Tesla) to 10 T in the examination region. Inalternative embodiments, the main magnet system may be implemented asconfiguration other than a solenoid and/or may be implemented as an openmagnet, such as vertical magnet or a double-donut magnet, and/or may beimplemented using lower fields (e.g., 0.1 T to 0.5 T) depending on thedesign and/or application. Typically, however, the direction of a lowmagnetic field can be oriented in a desired direction, for example,perpendicular to the patient bed (e.g., vertically), while the directionof a high field is usually horizontal.

As noted above, FIG. 1B schematically depicts in more detail the uppercross-sectional portion of the main magnet system shown in FIG. 1A. Asshown, vacuum chamber (housing) 116 encloses a vacuum space 132 thatsurrounds the main magnet and is evacuated to a vacuum of, for example,10⁻⁵ Torr or lower pressure (i.e., higher vacuum) by a vacuum system(not shown) comprising one or more vacuum pumps coupled to vacuum space132 via one or more ports, valves and/or feedthroughs, etc. Vacuumchamber housing 116 may be made of aluminum, stainless steel, or othermetallic or other non-metallic material, such as glass, ceramic,plastics, or combination of these materials. As will be understood bythose skilled in the art, vacuum space 132 provides thermal isolationbetween the cold main magnet and the room temperature wall of vacuumchamber housing 116.

The main magnetic coil 104, as well as the correction coils 106, may beimplemented as a low temperature superconductor (LTS) or as a hightemperature superconductor (HTS). A LTS main magnet may be made usingLTS wire, including, for example, NbTi, Nb₃Sn, Nb₃Al, MgB₂, and otherlow temperature superconductor wires. An HTS main magnet may be madeusing HTS tape, including, for example, one or more of YBCO, BSCCO, andother high-temperature superconductor tapes with critical temperatureabove 77K. As understood by those skilled in the art, one or more setsof correction coils 106 may be provided for the purpose of achievinggreater magnetic field uniformity. Such correction coils are typicallydesigned to carry only a small fraction of the current carried by themain superconductive coils, and/or have a small fraction of the numberof turns of the main superconductive coils, and the field contributionof a correction coil is designed to be nonuniform, so that incombination with the main magnetic field, the field of the correctioncoil acts to reduce overall magnetic field non-uniformity.

The superconducting magnet coils 104, as well as the superconductingcorrection coils 106, are wound onto main magnetic coil frame 108, whichmay be made from one or more materials such as stainless steel,aluminum, FR4 (e.g., self-extinguishing flammable G10), or othermechanically strong materials. The main magnetic coil frame is mountedin good thermal contact to thermal sink 110, which is thermally coupledto cryogenic system 160 such that heat is conducted from the mainmagnet, via thermal sink 110, to cryogenic system 160. Materialssuitable for making the thermal sink 110 include, for example, alumina,sapphire, and metal.

In some embodiments such as depicted in FIGS. 1A and 1B, cryosystem 160may be implemented as a two-stage system comprising a cryocooler 162, afirst stage 164, a second stage 168, wherein the first stage 164 isconnected to thermal shield 114 and the second stage is connected to thethermal sink 110 and/or to a cryogen, such as liquid helium, containedwithin cryogen container 112. In some embodiments, cryogen container 112may not be implemented, as cooling may be provided via thermal sink 110without using a surrounding cryogen. The temperatures of the first andsecond stage of the cryocooler are, for example, 40 K and 20 K,respectively, or 77 K and 40 K, respectively, or various othercombinations as desired, depending on various design parameters, such asthe material used for the superconducting magnet, the type of cryosystememployed, heat sources or loads, etc. Accordingly, thermal shield 114has a temperature between the room temperature vacuum wall and lowtemperature magnet coil, and thus, it will prevent radiation from theroom temperature vacuum wall from heating the superconductor mainmagnet. Some embodiments, however, may employ more than one layer ofthermal shielding or, alternatively, may not employ a thermal shield114.

In various embodiments, cryogenic system 160 may be implemented as anyof various single stage or multi-stage cryocoolers, such as, forexample, a Gifford McMahon (GM) cryocooler, a pulse tube (PT) cooler, aJoule-Thomson (JT) cooler, a Stirling cooler, or other cryocooler may

As shown in FIG. 1A, magnet/coil housing 130 also includes a secondvacuum chamber that comprises an interior portion of vacuum chamberhousing 116, an end-wall portion of magnet/coil housing 130, and aninterior wall 150, and that encloses a vacuum space 142, gradient coils103, and RF coil 105. The vacuum chamber enclosing vacuum space 142 iscoupled to a high vacuum pumping system to establish a low pressure(e.g., high vacuum condition) during manufacture, and is sealed after ahigh vacuum has been reached. RF coil 105 and gradient coils 103 areeach in thermal contact with a common heat sink 110, which is thermallycoupled to a cryogenic system 170 comprising a cryocooler 172 and asecond stage 174 having one end thermally coupled to the heat sink 110and its other end thermally coupled to the cryocooler 172. Illustrativematerials suitable for making the heat sink include ceramic such asalumina, crystals such as sapphire and metal, and glass.

In accordance with some embodiments of the present invention, a secondinterior wall 152 is provided to form another vacuum space 154, whichprovides additional thermal isolation, thus also enhancing user comfortwith respect to temperature. Vacuum space 154 may be coupled to a vacuumpump or may be implemented as a hermetically sealed chamber. The radialextent (i.e., with respect to cylindrical coordinates corresponding tothe generally cylindrical shape of the main magnet) of vacuum space 154is generally minimized so as to ensure that the RF coil 105 ismaintained close to the examination region. Illustrative materials forthe interior walls 150 and 152 include G10 fiberglass, glass, glasscomposites, or a combination of these materials. As known, thesematerials are non-magnetic and will not interfere with the gradientfields or RF signal in the examination region.

In this configuration, where the superconductive RF coil 105 and thesuperconductive gradient coils 103 are both commonly cooled, moretypically RF coil 105 and gradient coils 103 are implemented as the sametype of superconductor, namely, either HTS or LTS (although it isnevertheless possible to implement one of these elements as HTS and theother as LTS, provided they are cooled below the critical LTStemperature). A suitable form of an HTS RF coil and HTS gradient coilsfor this application is a superconductor tape made by, for example,Bismuth Strontium Copper Oxides (BSCCO). For example, detailed teachingof fabricating HTS RF coils from HTS tape is described in U.S. Pat. No.6,943,550, the disclosure of which is incorporated herein by reference.In alternative embodiments, the superconductor RF coil may beimplemented as a superconductor thin film, such as a superconductor thinfilm comprising an HTS material such as Yttrium Barium Copper Oxide(YBCO), Thallium-Barium-Calcium-Copper Oxide (TBCCO), MgB2, or MB,wherein M is selected from the group consisting of Be, Al, Nb, Mo, Ta,Ti, Hf, V, and Cr. Detailed teaching of fabricating HTS film coil on aflat substrate is described in Ma et al, “Superconducting MR SurfaceCoils for Human Imaging,” Proc. Mag. Res. Medicine, 1, 171 (1999) andthe disclosure of which is incorporated herein by reference in itsentirety. Additional teachings concerning HTS coils are described in Maet al., “Superconducting RF Coils for Clinical MR Imaging at Low Field,”Academic Radiology, vol. 10, no., 9, September 2003, pp. 978-987, and inMiller et al., “Performance of a High Temperature Superconducting Probefor In Vivo Microscopy at 2.0 T,” Magnetic Resonance in Medicine,41:72-79 (1999), the disclosures of which are incorporated herein byreference in their entirety.

As will be understood by those skilled in the art, RF coil 105 may beimplemented as separate coils for the RF transmitter and the RFreceiver, or as a common coil for both the transmitter and the receiver(i.e., a transceiver coil). Additionally, in some embodiments where thetransmitter and receiver coils are separate coils, only one of the coils(e.g., the receiver coil) may be implemented as a superconducting coil(e.g., the other coil may be implemented as a conventional copper coil).Additionally, in some embodiments, superconductive RF coil 105 may beimplemented as a coil array, such as an HTS coil array.

In some alternative embodiments of the present invention, one or more ofthe RF coils (e.g., the transmitter coil or the receiver coil, ifimplemented as separate coils) may be implemented as anon-superconducting coil that is formed from one or more materials thatwhen cooled to a given temperature (e.g., cryogenically cooled,refrigerated, water cooled, thermoelectrically cooled, etc.) has ahigher conductivity than that of copper at the given temperature. Suchnon-superconducting coils may be implemented, for example, fromsemiconductor two-dimensional electron gas (2DEG) material structures(e.g., GaAs and/or InP based), carbon nano-tubes, and other metals. Asused herein, for purposes of distinguishing between cryogenics andrefrigeration, a temperature approximately equal to or lower than about−73.3° C. (−100° F.) may be considered as being cryogenic.

Referring now to FIG. 2A, the gradient coils 103 of the illustrativesuperconductor MRI system of FIGS. 1A and 1B are depicted in more detailin an oblique view, in accordance with some embodiments of the presentinvention. In such embodiments, as depicted in FIG. 2A, threeindependent gradient coils for creating magnetic field variations alongthree orthogonal directions are formed or otherwise provided on and/orwithin the surfaces of three respective coaxial cylindrical supportstructures, namely, x-gradient support 258, y-gradient support 262, andz-gradient support 264. In accordance with typical convention, x- and y-indicate the two orthogonal directions perpendicular to the mainmagnetic field, and z- indicates the direction of the main magneticfield. Thus, the x-gradient support 258, y-gradient support 262, andz-gradient support 264 support respective gradient coils for providingmagnetic field gradients along the x-, y-, and z-directions,respectively. The gradient supports 258, 262, and 264 may be made of,for example, G10 or other non-ferromagnetic, non-conductive (e.g.,non-metallic, insulating) material. In this embodiment, the z-gradientcoil is a solenoid coil, and the x- and y-gradient coils are saddlecoils that each span or cover about half of their respective cylindricalsupports in the circumferential direction. The y-gradient support 262 ismounted in good thermal contact to x-gradient support 258 and toz-gradient support 264, which is mounted in good thermal contact tothermal sink 110. In various alternative embodiments, a heat sink may beadditionally or alternatively mounted in contact with x-gradient support258. When implemented in addition to thermal sink 110, such a heat sinkin contact with x-gradient support 258 may be cooled either bycryocooler 172 (i.e., the same cryocooler that cools thermal sink 110)or by a separate cryocooler. When implemented as an alternative tothermal sink 110 for cooling the gradient coils, thermal sink 110 may bethermally decoupled (e.g., spatially separated) from z-gradient support264, while still being thermally coupled to RF coil 105 for cooling theRF coil 105.

FIG. 2B schematically illustrates cylindrical x-gradient support 258 ofFIG. 2A depicted in a plan view, showing the x-gradient coil 268 that issupported by x-gradient support 258, in accordance with an embodiment ofthe present invention. The surface of the x-gradient support 258 isusually recessed (e.g., etched or carved) where the gradient coil 268(wire) is located, and the gradient coil wire is fixed and bonded in therecess so the wire will not move when current conducts through thegradient coil wire in the magnetic field (e.g., resulting in a Lorentzforce). The y-oriented gradient coil provided on y-gradient support 262has essentially the same design and construction as the x-orientedgradient coil 268 on x-gradient support 258, except for slightdimensional variations to account for the slightly smaller diameter ofthe y-gradient support compared to that of the x-oriented gradientsupport. The center 260 of the x-gradient coil 268 is facing thex-direction as indicated by FIGS. 2A and 2B, and the y-gradient coil isdisplaced 90° circumferentially relative to the x-gradient coil. Thesolenoidal z-gradient coil (not shown in detail) is similarly fabricatedon and/or within the surface of the z-gradient support 264, but with thez-gradient coil wound helically about the cylindrical axis of z-gradientsupport 264, with half of the coil along the cylindrical axis wound inthe same direction as the main magnet winding such that the z-gradientcoil increases the magnetic field within this half of the coil, and withthe other half of the coil along the cylindrical axis wound in theopposite direction such that the z-gradient coil decreases the magneticfiled within this other half of the coil.

In some embodiments, as further discussed hereinbelow, gradient coils103 and RF coil 105 may be separately cooled and thermally isolated fromeach other, which may be desirable, for example, to provide differentoperating temperatures for the gradient coils and the RF coil (e.g.,when different materials are used for these elements). Such alternativeembodiments may include disposing the RF coils and the gradient coilseither in a common vacuum chamber or in separate vacuum chambers.

In some embodiments, such as the embodiments discussed above inconnection with FIGS. 1A and 1B, the main magnet, the gradient coils,and the RF coil are all implemented as superconductors, and each ofthese components may be implemented as either HTS or LTS, thus providingfor eight (8) possible permutations, assuming all of the gradient coilsare implemented with the same type of superconductor (i.e., either HTSor LTS). In accordance with some preferred embodiments of the presentinvention, the main magnet, the gradient coils, and the RF coil are eachimplemented with HTS materials. As will be appreciated by those skilledin the art, such an all-HTS configuration provides many advantages interms of providing for a cost-effective, high quality, high performanceMRI system.

For instance, superconducting main magnets made from low temperaturesuperconductors are generally very bulky and heavy. A main magnet madeof HTS in accordance with various embodiments of the present invention,however, is comparatively much lighter and more compact as, for example,the same magnetic field magnitude may be achieved with less HTS wirethan LTS wire. Additionally, because it can be operated at a much highertemperature (e.g., 77K) than an LTS magnet (e.g., around 10-20 K), anHTS main magnet uses much less cryogen and hence reduces costsubstantially. Similarly, implementing both the gradient and RF coilswith HTS materials also reduces cooling costs while also simplifyingthermal and vacuum isolation design compared to embodiments of thepresent invention that employ LTS materials for the gradient coilsand/or RF coils. At the same time, compared to conventional copper RFcoils and gradient coils, overall MRI system performance issignificantly enhanced due, in part, to the HTS RF coils providing forhigh sensitivity (e.g., reduced coil noise and hence higher SNR), whilethe HTS gradient coils provide for high drive currents, rapid switching,and significantly reduced heat dissipation.

Referring now to FIGS. 3A-3D, schematically depicted are differentexamples of cooling configurations that may be used within an MRI systemaccording to various embodiments of the present invention. As shown inFIG. 3A, each of the superconducting coils 202 are individually cooledwithin their own cooling chamber by a separate cryogenic cooling system204. Main magnet coils 206 are cooled to exhibit HTS or LTScharacteristics under the control of cryogenic cooling system 208.Similarly, gradient coils 210 are cooled to exhibit HTS or LTScharacteristics under the control of cryogenic cooling system 212. Also,RF coils 214 are cooled to exhibit HTS or LTS characteristics under thecontrol of cryogenic cooling system 216.

As shown in FIG. 3B, the main magnet coils 220 are cooled to exhibit HTSor LTS characteristics under the control of cryogenic cooling system222. However, gradient coils 226 and RF coils 228 are cooled to exhibitHTS or LTS characteristics under the control of common cryogenic coolingsystem 230. Within this embodiment, the coils are all cooled withintheir own individual cooling chamber.

As shown in FIG. 3C, the main magnet coils 234 are cooled to exhibit HTSor LTS characteristics under the control of cryogenic cooling system236, whereby coils 234 are cooled within their own cooling chamber.However, both the gradient and RF coils 238 are cooled to exhibit HTS orLTS characteristics under the control of common cryogenic cooling system240. With this embodiment, both the gradient and RF coils 238 are cooledwithin the same cooling chamber.

As shown in FIG. 3D, the main magnet and gradient coils 244 are bothcooled to exhibit HTS or LTS characteristics under the control ofindividual cryogenic cooling system 246, whereby coils 244 are bothcooled within the same cooling chamber. The RF coils 248, however, arecooled to exhibit HTS or LTS characteristics under the control ofindividual cryogenic cooling system 250, whereby the RF coils 248 arecooled within a separate cooling chamber to that of the main magnet andgradient coils 244.

Additionally, as will be understood by those skilled in the art in viewof the foregoing, various embodiments of the present invention may beimplemented with the main magnet, gradient coils, and RF coil beingcooled by a common cryocooler, regardless of whether the main magnet,gradient coils, and RF coils are each disposed in separate (respective)vacuum isolated cooling chambers, or are disposed in two vacuum isolatedcooling chambers (e.g., gradient coils and RF coil(s) in same chamber),or are disposed in a common vacuum isolated cooling chamber.

FIG. 4A illustrates a first cross sectional view of an illustrative coilconfiguration 300 associated with a superconducting MRI system employinga cylindrical, solenoid main magnet structure (e.g., similar to MRIsystem 100 shown in FIG. 1A) according to some embodiments. Theconfiguration 300 includes a first vacuum chamber 316, a second vacuumchamber 314, one or more main magnet coils 302, one or more gradientcoils 304, one or more RF coils 306, and walls 308, 310, and 312. Aswill be understood in view of the further description below, inaccordance with various embodiments, each of one or more of walls 308,310, and 312 in configuration 300 may be implemented as a hermeticallysealed double-walled structure, which, in some embodiments, may beimplemented in accordance with, or similar to, the hermetically sealeddouble-walled structures (and vacuum thermal isolation housing)described in U.S. application Ser. No. 12/212,122, filed Sep. 17, 2008,and in U.S. application Ser. No. 12/212,147, filed Sep. 17, 2008, eachof which is herein incorporated by reference in its entirety.

The first vacuum chamber 316 houses the super MRI magnet and itscorresponding main magnet coil 302. Vacuum chamber 316 is formed betweenhermetically sealed double-walls 308 and 310, whereby the cavity withineach of double-walls 308 and 310 is vacuum pumped, filled (optionally)with thermal insulation material (e.g., fiber glass), and appropriatelysealed (e.g., via melding) to maintain a high-grade vacuum. Theenclosure associated with the first vacuum chamber 316 is also evacuatedusing a suitable vacuum pump. The outer double-wall 308 of the firstvacuum chamber 316 may be constructed from conventional vacuum chambermaterials, such as, but not limited to, aluminum or stainless steel. Theinner double-wall 310 of the first vacuum chamber 316 may, however, beproduced from a non-magnetic and non-metallic material, such as, butlimited to, glass, non-conductive ceramic, G10, FR4, or plastic.

As previously described, once a sufficient vacuum is created within thefirst vacuum chamber 316, a cryogenic cooling system is used to reducethe temperature of the main magnet coil 302. The required temperaturereduction may depend on the coil material. By utilizing either lowtemperature superconducting (LTS) material or high temperaturesuperconducting (HTS) materials in the construction of coil 302, itsresistance is greatly reduced in comparison to conventionally cooledcopper coils. The superconducting windings of main magnet coil 302 will,therefore, reduce the amount of heat generation/dissipation that occurswithin the coil windings when driven by an established current necessaryto generate a particular target magnetic field (e.g., 1 Tesla). Also, asa consequence, the amount of power required to generate and maintain theparticular magnetic field by the main MRI magnet is reduced. Moreover,future MRI applications may lead to the use of higher magnetic fieldmagnitudes (e.g., greater than 7 Tesla). Under such circumstances, theuse of superconductive main magnet coils enables the generation ofhigher current densities in the coil and thus, increased magnetic fieldcapabilities. The cryogenic cooling system may, for example, operateover a range of 20-40 Kelvin (K). Also, according to some embodiments, asuperconducting main magnet coil may have a length of 0.5-3 meters (m),an outer diameter of 1-3 m, an inner diameter of 0.1-2.5 m, and asubstantially cylindrical geometry.

The second vacuum chamber 314 houses both the gradient coils 304 and theRF coils 306. Vacuum chamber 314 is formed between hermetically sealeddouble-walls 310 and 312, whereby the cavity within each of double-walls310 and 312 is also vacuum pumped, filled (optionally) with thermalinsulation material (e.g., fiber glass), and appropriately sealed (e.g.,via melding) to maintain a high-grade vacuum. The enclosure associatedwith the second vacuum chamber 314 is also evacuated using a suitablevacuum pump. The outer double-wall 310 of the second vacuum chamber 314is produced from a non-magnetic and non-metallic material, such as, butlimited to, glass, non-conductive ceramic, G10, FR4, or plastic. Theinner double-wall 312 of the second vacuum chamber 314 is, however,materially constructed to have no screening effect on RF signalstransmitted by and received from the RF coils 306, and produces no eddycurrent effects that may result from the application of gradient signalsto the gradient coils 304.

Once a sufficient vacuum is created within the second vacuum chamber314, another cryogenic cooling system is used to reduce the temperatureof either or both the gradient coils 304 and RF coils 306. As previouslymentioned, the required temperature reduction may depend on the coilmaterial. By utilizing either low temperature superconducting (LTS)material or high temperature superconducting (HTS) materials in theconstruction of coils 304 and/or 306, there respective resistances aregreatly reduced in comparison to conventionally cooled copper coils orother such non-superconducting materials. The superconducting windingsof gradient coils 304 (LTS or HTS) minimizes/reduces the amount ofgradient heating, and allows for rapid switching of high gradientfields. Thus, faster image acquisition (increased temporal resolution)and a reduction in additional cooling requirements for dissipatinggradient coil generated heat are realized. The cryogenic cooling systemassociated with cooling the gradient coils 304 may, for example, operateover a range of 40-60 Kelvin (K). According to some embodiments, asuperconducting gradient coil may include a length of 0.2-2 meters (m),an outer diameter of 0.1-2.5 m, an inner diameter of 0.02-2.3 m, and acylindrical solenoid and saddle geometry. A superconducting RF coil(HTS) may include a length of 0.01-0.5 m, an outer diameter of 0.02-1.0m, an inner diameter of 0.01-0.8 m, and a cylindrical solenoid andsaddle geometry. The superconducting RF coils 306 reduce the coil noise.This in turn results in an increased S/N performance within the RFreceiver circuitry (provided that the sample noise does not overwhelmthe coil noise), which provides for faster acquisition and/or improvedimage resolution capture. The cryogenic cooling system associated withcooling the gradient coils 304 and RF coils may, for example, operateover a range of 40-60 Kelvin (K). FIG. 4B illustrates a second crosssectional view of the exemplary coil configuration 300 taken along alongitudinal direction.

Many different HTS and LTS materials may be employed in the constructionand operation of the superconducting MRI system. For example, thegradient coils 304 may be constructed from Bi-223 tape, which is acommercial low-cost HTS material. In some instances, the Bi-223 tape maybe sheathed by pure silver (Ag) in order to enhance its mechanicalstrength. When the Bi-223 tape is cooled by, for example, immersion inliquid nitrogen, it exhibits superconducting properties, whereby itsresistance reduces to approximately zero. The superconducting RF coils306, which can be configured as either a transceiver or discretetransmitter and receiver, may also be formed from HTS materials (e.g.,YBaCuO, BiSrCaCuO, etc.), as well as other superconductors,nano-materials such as carbon nano-tubes, and two-dimensional electrongas (2DEG) materials/structures having high conductivity characteristics(e.g., using GaAs or InP material system based compounds). Alternativelyor additionally, the HTS RF coils 306 may include an array of thin filmcoils, each having a substrate diameter of, for example, about 1 cm to30 cm. The superconducting main magnet coils may be constructed fromeither HTS or LTS materials. For example, LTS materials such as MgB₂(Magnesium Di-boride) may be used to form the main magnet coil 302. Froma cooling perspective, the superconducting coils may be operated overdifferent temperature ranges. For example, the superconducting mainmagnet coils 302 may be cooled over a range of about 20-40K. Thesuperconducting gradient coils 304 may be cooled over a range of about40-60 K, while the superconducting RF coils 306 may be maintained athigher temperatures ranging from about 40-60 K or at about 77K.Alternatively, both the superconducting gradient coils 304 andsuperconducting RF coils 306 may be cooled to around 77K, while thesuperconducting gradient coils 304 is cooled over a range of 20-40K. Amyriad of different operating temperatures may be used. For example,according to some configurations, all the superconducting coils may bemaintained around a temperature of 77K.

Although the described embodiments show the coils configured in a mannerthat provides a horizontal magnetic field, other MRI systems mayincorporate structural designs that facilitate the generation ofvertical magnetic fields of differing strength (e.g., 0.5 T. 1.0 T,etc.) across various fields of views (FOV). Such MRI system examplesinclude, but are not limited to, an asymmetric head-scanning MRIincorporating a 6 or 8 RF coil array; an orthopedic MRI system (0.2-0.5T system using a Helmholtz Coil Pair) for examination of hands of legs;or an open vertical field MRI system for scanning breasts, whereby theRF coils may be built into the examination bed. The open vertical fieldMRI system design concept may also be extended for examining animals. Itmay also be appreciated that while the MRI system embodiments describedhereinabove are typically directed to detecting hydrogen atoms withinthe water of bodily tissue, it may be adapted to detect other nuclei.

The present invention has been illustrated and described with respect tospecific embodiments thereof, which embodiments are merely illustrativeof the principles of the invention and are not intended to be exclusiveor otherwise limiting embodiments. Accordingly, although the abovedescription of illustrative embodiments of the present invention, aswell as various illustrative modifications and features thereof,provides many specificities, these enabling details should not beconstrued as limiting the scope of the invention, and it will be readilyunderstood by those persons skilled in the art that the presentinvention is susceptible to many modifications, adaptations, variations,omissions, additions, and equivalent implementations without departingfrom this scope and without diminishing its attendant advantages. Forinstance, except to the extent necessary or inherent in the processesthemselves, no particular order to steps or stages of methods orprocesses described in this disclosure, including the figures, isimplied. In many cases the order of process steps may be varied, andvarious illustrative steps may be combined, altered, or omitted, withoutchanging the purpose, effect or import of the methods described. It isfurther noted that the terms and expressions have been used as terms ofdescription and not terms of limitation. There is no intention to usethe terms or expressions to exclude any equivalents of features shownand described or portions thereof. Additionally, the present inventionmay be practiced without necessarily providing one or more of theadvantages described herein or otherwise understood in view of thedisclosure and/or that may be realized in some embodiments thereof. Itis therefore intended that the present invention is not limited to thedisclosed embodiments but should be defined in accordance with theclaims that follow.

What is claimed is:
 1. A system configured for magnetic resonanceimaging (MRI) and/or magnetic resonance spectroscopy, the systemcomprising: a superconducting main magnet operable to generate a uniformmagnetic field in an examination region, wherein the superconductingmain magnet is a cylindrical solenoid magnet comprising windings thatextend over a longitudinal axis and surround a bore that is disposedinterior thereto and that comprises said examination region; at leastone superconducting gradient field coil that is disposed between saidexamination region and said superconducting main magnet, and that isoperable to apply a respective at least one magnetic field gradientwithin the examination region; and at least one superconducting RF coilthat is disposed between said examination region and saidsuperconducting main magnet, and that is operable to transmit andreceive radio frequency signals to and from the examination region;wherein said superconducting main magnet, each of said at least onesuperconducting gradient field coil, and each of the at least onesuperconducting RF coil, each comprise a high temperaturesuperconductive (HTS) material.
 2. The system according to claim 1,wherein the same HTS material is used for said superconducting mainmagnet, each of said at least one superconducting gradient field coil,and each of the at least one superconducting RF coil.
 3. The systemaccording to claim 2, wherein the HTS material comprises bismuthstrontium copper oxide (BSCCO) formed as a tape.
 4. The system accordingto claim 1, wherein said at least one gradient coil and the at least oneRF coil are disposed in at least one vacuum chamber having at least onenon-magnetic and non-metallic wall disposed between the examinationregion and the gradient coil and the at least one RF coil, and whereinthe at least one gradient coil and the at least one RF coil are disposedin a common vacuum chamber comprising said at least one non-magnetic andnon-metallic wall.
 5. The system according to claim 4, furthercomprising a further vacuum chamber disposed between said common vacuumchamber and the examination region, said further vacuum chambercomprising a first wall formed from said at least one non-magnetic andnon-metallic wall, and a second non-magnetic and non-metallic wallspaced away from said first wall.
 6. The system according to claim 1,wherein said at least one gradient coil and the at least one RF coil aredisposed in at least one vacuum chamber having at least one non-magneticand non-metallic wall disposed between the examination region and thegradient coil and the at least one RF coil, wherein said at least onevacuum chamber comprises a first vacuum chamber containing said at leastone gradient coil, and a second vacuum chamber disposed between thefirst vacuum chamber and the examination region and containing said atleast one RF coil, and wherein said at least one non-magnetic andnon-metallic wall comprises a first non-magnetic and non-metallic walldisposed between the examination region and the gradient coil, and asecond non-magnetic and non-metallic wall disposed between theexamination region and the at least one RF coil.
 7. The system accordingto claim 6, further comprising a further vacuum chamber disposed betweensaid second vacuum chamber and the examination region, said furthervacuum chamber comprising a first wall formed from said secondnon-magnetic and non-metallic wall, and a third non-magnetic andnon-metallic wall spaced away from said first wall.
 8. The systemaccording to claim 1, wherein said main magnet is disposed in a firstvacuum chamber, and said at least one RF coil and said at least onegradient coil are disposed in a second vacuum chamber.
 9. The systemaccording to claim 1, wherein said main magnet, said at least one RFcoil, and said at least one gradient coil are disposed in respectivevacuum chambers.
 10. The system according to claim 1, wherein said atleast one RF coil and said at least one gradient coil are disposed in acommon vacuum chamber.
 11. The system according to claim 10, whereinsaid at least one RF coil and said at least one gradient coil arethermally coupled to a common heat sink.
 12. The system according toclaim 1, wherein said at least one RF coil comprises an RF coil array.13. The system according to claim 1, wherein the at least one RF coilcomprises a single RF coil operable as both a transmitter and areceiver.
 14. The system according to claim 1, wherein the at least oneRF coil comprises a transmitter RF coil and a receiver RF coil.
 15. Thesystem according to claim 1, wherein said superconducting main magnet isconfigured for cooling by a first cryogenic cooling system, said atleast one RF coil is configured for cooling by a second cryogeniccooling system, and said at least one gradient coil is configured forcooling by a third cryogenic cooling system, wherein the first, second,and third cryogenic cooling systems are part of the system configuredfor MRI and/or magnetic resonance spectroscopy.
 16. The system accordingto claim 1, wherein said superconducting main magnet is configured forcooling by a first cryogenic cooling system, and said at least one RFcoil and said at least one gradient coil are configured for cooling by asecond cryogenic cooling system, wherein the first and second cryogeniccooling systems are part of the system configured for MRI and/ormagnetic resonance spectroscopy.
 17. The system according to claim 1,wherein said at least one superconducting gradient field coil comprisesthree superconducting gradient field coils that are configured toprovide magnetic field gradients in three respective orthogonaldirections, wherein one of the directions is along the direction of theuniform magnetic field generated in the examination region by thesuperconducting main magnet.